The present invention relates to the magnetic resonance arts. It finds particular application in conjunction with medical magnetic resonance imaging systems and will be described with particular reference thereto. It is to be appreciated, however, that the invention will also find application in conjunction with other types of magnetic resonance imaging systems, magnetic resonance spectroscopy systems, and the like.
In magnetic resonance imaging, a strong uniform static magnetic field B.sub.0 is generated, often by a superconducting magnet. This static magnetic field B.sub.0 polarizes the nuclear spin system of an object to be imaged. Superconducting magnets are commonly wound on a cylindrical body former mounted in an annular helium vessel surrounded by an annular vacuum vessel for thermal isolation. The superconducting magnet generates the static magnetic field, B.sub.0 along its own longitudinal axis and the common longitudinal axis of the cylindrical bore of the vacuum vessel, commonly denoted as the "z-axis".
To generate a magnetic resonance signal, the polarized spin system is first excited by a radio frequency magnetic field perpendicular to the z-axis. This RF field, denoted B.sub.1, is produced by an RF coil located inside the bore of the magnet and closely conforming thereto to maximize the space available to receive a patient. The RF magnetic field, which is changing in time in a sinusoidal waveform, is turned on and off to create short RF pulses to excite magnetization in the polarized object in the bore. More specifically, the RF pulses tip the magnetization out of alignment with the z-axis and cause its macroscopic magnetic moment vector to precess around the z-axis. The precessing magnetic moment, in turn, generates a radio frequency magnetic resonance signal that is received by the RF coil in a reception mode.
To encode a sample spatially, magnetic field gradient pulses are applied after the RF excitation. The gradient magnetic fields are also applied in pulses to generate magnetic fields pointing in the z-axis, but changing in magnitude linearly in x, y, or z-directions. These gradient pulses are commonly denoted as G.sub.x, G.sub.y, and G.sub.z pulses, respectively. The gradient magnetic fields are generated by gradient magnetic field coils which are also located inside the magnet bore. Commonly, the gradient field coils are mounted on a cylindrical former between the RF coil and the bore. The RF and gradient field coils have a sufficient inner diameter to receive the entire body of a patient within their circular bore.
These cylindrical whole body RF and gradient coils are routinely used to examine the patient's chest cavity and to examine individual organs therein, such as the heart. Examining individual organs or small regions of the chest cavity places the RF coils a significant distance from the region of interest.
The most common whole body RF coils are saddle coils and birdcage coils. Both saddle and birdcage coils include electrical conductors which are mounted on a cylindrical dielectric former. Prior patents have also suggested mounting the RF coils on an elliptical former to follow the contour of a patient's body more accurately. Typical saddle and birdcage coils are described in "Radio Frequency Coils" by Hayes, Edelstein, and Schenck, NMR in Medicine: The Instrumentation and Clinical Applications, pages 142-165, (1985). Typical saddle coils includes four conductors extending in the z-direction, adjacent pairs of which are connected at their ends by connectors extending along the cylindrical former. In a birdcage coil, a plurality of conductors, e.g. eight, are spaced equidistant along the circular former and interconnected at their ends by a generally annular conductor extending along the surface of the circular former. For quadrature detection, a second coil rotated 90.degree. from the first is provided.
In order to accommodate the entire body of a patient, conventional birdcage and saddle coils are relatively large in diameter and length, e.g. 60 cm in diameter and about the same length. One of the problems with whole body birdcage and saddle coils is that the large size causes a loss of signal to noise ratio. The signal to noise ratio is generally proportional to coil sensitivity and to the coil filling factor. The coil sensitivity is generally inversely proportional to a power of the coil radius. Thus, the larger the diameter of the coil, the poorer its sensitivity and the lower its signal to noise ratio. The filling factor is the ratio of the magnetic energy stored in the region of interest in the patient and the total energy generated by the RF coil. When viewing only a small region of the patient, such as the heart, the magnetic energy stored in the heart region is relatively small compared to the total energy generated by the RF coil. Hence, the filling factor is low and the signal to noise ratio is also low.
In order to improve the signal to noise ratio, when examining individual organs, the whole body RF coil is often used only as a transmitter coil. A separate receiver coil is positioned on the surface of the patient to detect the MR signals. The dedicated receiver coils are each designed for a specific organ and are sized and shaped accordingly. This increases its sensitivity to the organ and suppresses noise from other tissue in the patient. Thus, a large inventory of coils is required to image various internal organs of the body. Moreover, the sensitivity of surface coils falls off sharply away from the plane of the coil. This drop off in sensitivity toward the deeper regions of the subject causes non-uniformity in the images. Moreover, the sensitivity may drop off sufficiently fast that the examination of some deeper organs is impractical.
The present invention provides a new and improved RF coil which overcomes the above-referenced problems and others.